Since the introduction of the first implantable pacemakers in the early 1960s, there have been considerable advancements both in the field of electronics and in the field of medicine, such that there is presently a wide assortment of commercially-available implantable medical devices. The class of implantable medical devices now includes not only pacemakers, but also implantable cardioverters, defibrillators, neural stimulators, and the like. Today's state-of-the-art implantable medical devices are vastly more sophisticated and complex than early pacemakers, and are capable of performing significantly more complex functions. The therapeutic benefits of implantable medical devices have been well proven.
An early pacemaker was disclosed in U.S. Pat. No. 3,057,356 issue to Greatbatch in 1962 and entitled "Medical Cardiac Pacemaker". The Greatbatch pacemaker included a relaxation oscillator for controlling the pacemaker to generate electrical cardiac stimulating pulses. Thus, the pacemaker operated asynchronously to provide fixed-rate cardiac stimulation not automatically changed in accordance with the patient's needs. The Greatbatch pacemaker proved to be effective in alleviating the symptoms of complete heart block. As an asynchronous device, however, the Greatbatch pacemaker had the possible disadvantage of operating to compete with the natural, physiological functioning of the heart during episodes of normal sinus condition.
Since 1962, implantable pulse-generating medical devices have been continuously evolving. For example, in order to overcome the possible disadvantages with asynchronous pacemakers, implantable pacemakers of the synchronous or demand type were developed wherein stimulating pulses are delivered only when required, and are not delivered when the heart functions with a normal sinus rhythm. An early demand-type pacemaker is disclosed, for example, in U.S. Pat. No. 3,478,746 entitled "Cardiac Implantable Demand Pacemaker". The demand pacemaker solves the problem arising with asynchronous pacemakers by inhibiting delivery of stimulating pulses in the presence of detected ventricular activity, and by delivering stimulating pulses only in the absence of natural cardiac activity.
Another improvement which occurred since the first implantable cardiac pacemaker is the ability to reprogram certain operational parameters of the pacemaker after it has been implanted. For example, in U.S. Pat. No. 3,805,796 issued to Terry, Jr. et al. in 1974 and entitled "Implantable Cardiac Pacemaker Having Adjustable Operating Parameters". The Terry, Jr. device included circuitry to allow the rate of the pacemaker to be non-invasively changed after the device was implanted. The stimulation rate was varied according to the number of times that a magnetically actuated reed switch was closed. The device operated to count the number of times the reed switch was closed and storing that count in a binary counter. Each state of the counter was connected to either engage or bypass one resistor in a serially-connected resistor chain, where the resistor chain formed part of the RC time constant controlling pacemaker rate.
The concept of the Terry, Jr. patent has also been improved upon, as exemplified in U.S. Pat. No. 4,066,086 to Adams et al. entitled "Programmable Body Stimulator" The Adams et al. patent discloses a pacemaker that responds to the application of radio frequency (RF) bursts while a magnetic filed held in close proximity to a reed switch in the device holds the reed switch closed. In the adams et al. circuit, only the rate is programmable in response to the number of RF bursts applied. The use of RF signals to program cardiac pacemakers was earlier disclosed in U.S. Pat. No. 3,833,005 issued to Wingrove in 1974 and entitled "Compared Count Digitally Controlled Pacemaker". The Wingrove device was capable of having both pacing rate and pacing pulse width programmed.
Perhaps the most significant advance in implantable device technology, however, was the incorporation of digital circuitry in implantable devices. Implantable device technology initially lagged behind conventional state-of-the-art electronic technology in its utilization of digital circuits. A primary reason for the delay in was that early digital circuits consumed unacceptably large amounts of energy to be used in battery-powered implantable devices impractical. Of course, conservation of battery power in implantable devices has always been a major concern in pacemaker design. Thus, although there were suggestions in the art to utilized digital techniques in cardiac pacemakers even as early as 1966 (see, e.g., Walsh et al., "Digital Timing Unit for Programming Biological Stimulators", American Journal of Medical Electronics, First Quarter, 1977, pp. 29-34), the first patent suggesting digital techniques in the context of cardiac pacemakers was U.S. Pat. No. 3,557,796 issued to Keller, Jr., et al. in 1971 and entitled "Digital Counter Driven Pacer".
The Keller, Jr. pacemaker included an oscillator driving a binary counter. When the counter reached a certain value, a signal was provided which caused a cardiac stimulating pulse to be provided. At the same time, the counter was reset and began counting oscillator pulses. The Keller, Jr. pacemaker also incorporated a demand feature, wherein the counter was reset upon detection of a natural heartbeat, as well as a refractory feature, wherein output pulses were inhibited for a certain time after the provision of a cardiac stimulating pulse or natural beat.
Improvements in digital technology and in battery technology have been such that the use of digital circuitry in implantable devices has, over the years, become increasingly feasible and increasingly common. Patents disclosing digital techniques useful in cardiac pacemakers include U.S. Pat. No. 3,631,860 to Lopin entitled "Variable Rate Pacemaker"; U.S. Pat. No. 3,857,399 to Zacouto entitled "Heart Pacer"; U.S. Pat. No. 3,865,119 to Svensson et al. entitled "Heartbeat Accentuated with Controlled Pulse Amplitude"; U.S. Pat. No. 3,870,050 to Greatbatch entitled "Demand Pacer"; U.S. Pat. No. 4,038,991 to Walters entitled "Cardiac Pacer with Rate Limiting Means"; U.S. Pat. No. 4,043,347 to Renirie entitled "Multiple-Function Demand Pacer with Low Current Drain"; U.S. Pat. No. 4,049,003 to Walters et al. entitled "Digital Cardiac Pacer"; and U.S. Pat. No. 4,049,004 to Walters entitled "Implantable Digital Cardiac Pacer Having Externally Selectable Operating Parameters and One-Shot Digital Pulse Generator for Use Therein".
Pacemakers incorporating digital circuitry are also disclosed in U.S. Pat. No. 4,250,883 issued to David L. Thompson and entitled "Digital Cardiac Pacemaker"; and in U.S. Pat. No. 5,052,388 to Sivula et al. entitled "Method and Apparatus for Implementing Activity Sensing in a Pulse Generator". The Thompson '883 and Sivula et al. '388 patents are hereby incorporated by reference herein in their respective entireties.
The accuracy and reliability of digital circuits are factors that have encouraged their use in implantable devices. Their ability to be programmed and reprogrammed to alter one or more operating parameters further enhances their utility. For example, the pacemaker disclosed in the above-referenced Sivula et al. patent respond to radio frequency signals from a microprocessor-based external programming unit to alter numerous operating parameters, including pulse rate, pulse width and/or pulse amplitude, pacing mode, sensing mode and sensitivity, activity/rate response settings, refractory periods, AV-delay settings, and others. In U.S. Pat. No. 4,340,062 to Thompson et al. entitled "Body Stimulator Having Selectable Stimulation Energy Levels", there is disclosed a pacemaker in which the amplitude, duration, and repetition rate of cardiac stimulating pulses is externally controllable. The Thompson '062 patent is hereby incorporated by reference herein in its entirety.
Since digital technology has made it possible to provide numerous non-invasively programmable parameters in implantable devices, it is now relatively common for pacemakers to provide for a plurality of different stimulating pulse amplitude settings. One reason for the desirability of having programmable pulse amplitude in a pacemaker is that battery longevity can be maximized through selection of a pacing pulse amplitude appropriate for a given patient's pacing threshold. That is, for a patient with a relatively low pacing threshold, the pacing amplitude can be set to a correspondingly lower level than for a patient with a higher pacing threshold, thereby minimizing power consumption while at the same time ensuring that the pacing pulses will be sufficient to capture the patient's heart.
One difficulty in implementing programmable pulse amplitude in a pacemaker is ensuring that pacing pulses will be delivered at the selected pacing amplitude throughout the life of the pacemaker, even though the battery voltage will not remain at the same level at all battery depletion levels. Typically, implantable pulse generator utilize output capacitors to store the energy for an output pulse. Charging circuitry is provided to couple the output capacitors to the battery prior to delivery of a pulse. The charge accumulated on the output capacitors can be controlled, for example, by controlling the amount of time that the output capacitors are coupled to the battery voltage. However, such an arrangement assumes that the battery voltage remains the same throughout its life, the output capacitors always charging to the same voltage for a given charging interval. This, of course, is not a valid assumption.
To overcome the problem of declining battery voltage with battery discharge, a pacemaker may be provided with circuitry for controlling the process of charging of the output capacitors. For example, when an output capacitor is to be charged to a chosen voltage, the output capacitors may be coupled to a battery and to a comparator circuit. The charging circuit can then operate to keep the output capacitors coupled to the battery until the comparator circuit indicates that the desired voltage level has been reached, at which time the charging of the capacitors is discontinued. This arrangement does not assume a constant battery voltage for all depths of battery discharge, but has the drawback of increasing the complexity of the charging circuitry. Also, the regulation circuitry itself consumes power and thus increases the device's overall current consumption and reduces its projected longevity.
Lithium-iodine batteries are among the most commonly used power sources for modern implantable devices, and much has come to be known about their depletion characteristics. In particular, it is well known in the art that the output voltage from lithium-iodine batteries is relatively linear during early stages of depletion, but drops off rather sharply before end-of-life (EOL). This is due in part to the internal resistance of lithium-iodine batteries, which is relatively linear as a function of energy depletion until near EOL, at which time the resistance curve exhibits a "knee" where internal resistance begins to rise rapidly.
In typical lithium-iodine batteries, the cell cathode consists of molecular iodine weakly bonded to polyvinyl pyridine (P2VP). The initial cathode composition of lithium-iodine batteries is often expressed as the weight ratio of l.sub.2 to P2VP. Typical values of this ratio range from 20:1 to 50 1. No electrolyte as such is included in the construction of the cell, but a lithium iodine (Lil) electrolyte layer forms during cell discharge, between the anode and the cathode. The Lil layer presents an effective internal resistance to Li+ions which travel through it. Since the Lil layer grows with the charge drawn from the battery, this component of the battery resistance increases linearly as a function of energy depletion. In the implantable device context, where there is typically a relatively continuous energy depletion, this component of the internal resistance increases continually over time. However, particularly for a demand type pacemaker which at any given time may or may not be called upon to deliver stimulating pulses, the increase in this component is continuous but not necessarily linear with time, due to the fact that current drain is not constant.
Another component of internal resistance in lithium-iodine cells is caused by depletion of iodine in the cathode. The cathode is essentially a charge transfer complex of iodine and P2VP, and during discharge of the cell iodine is extracted from this complex. As noted above, the weight ratio of l.sub.2 to P2VP at beginning of life may range from 20:1 to 50:1. During extraction of iodine from the complex, the resistance to this process is low until the point is reached where the l.sub.2 -to-P2VP ratio is reduced to approximately 8:1, the ratio at which the cathode becomes a single phase and the iodine activity begins to be less than unity. At this point the resistance rises sharply. This gives rise to a non-linear internal resistance component which, for the lithium-iodine cell, is called variously the depletion resistance, depolarizer resistance, the charge-transfer complex resistance, or the pyridine resistance. By whatever names, the combination of the non-linear component with the linear component produces an overall resistance curve with a knee occurring toward EOL, the knee being caused by the reaching of the depletion of available charge carriers from the cathode.
Since it is often extremely critical for patients' well-being that implantable devices do not cease operating, it is common for implantable devices to monitor the level of battery depletion and to provide some indication when the depletion reaches a level at which the battery should be replaced. Pacemakers manufactured by Medtronic, Inc., for example, typically provide, for example via telemetry, an "elective replacement indictor" (ERI) when the battery depletion reaches a level such that replacement will soon be needed. Pacemakers may also provide an indication when the level battery depletion is such that the device must be replaced immediately. Other pacemakers may provide information about battery depletion levels throughout the device's life, for example, whenever the pacemaker is interrogated by an external programmer.
In the prior art, some ERI arrangements in implantable devices evaluate battery life based simply upon the terminal voltage of the battery, indicating ERI or EOL when the voltage falls below a predetermined threshold. However, due to the internal impedance characteristics of the battery, discussed above, terminal voltage can vary significantly depending upon current consumption. Thus, if relatively little current is drawn from the battery for a period of time when the battery is nearing but has not reached the ERI point, a sudden prolonged period of high demand on the battery may cause a situation in which too little time is available between ERI and total depletion of the battery. For a particular pacemaker and electrode combination in a given patient, there will be a variation in the effective load on the lithium-iodine battery, and a resulting variation in the overall current drain. Accordingly, if ERI is predicated upon sensing the voltage of the battery and detecting when it drops below a certain level, there can be very little assurance that the level chosen will correspond to the knee of the internal resistance curve.
It has been recognized in the prior art that since remaining battery life is directly related to the internal impedance of the battery itself, remaining battery life can be reliably predicted through accurate measurement of internal battery impedance. In U.S. Pat. No. 5,137,020 issued to Wayne et al. and assigned to the assignee of the present invention, there is described a battery impedance measuring arrangement wherein a current source and a reference impedance are applied to a battery which has been isolated from the remainder of the pacemaker circuitry. The Wayne et al. '020 patent is hereby incorporated by reference in its entirety into the present disclosure.
Other battery impedance measuring arrangements are proposed, for example, in U.S. Pat. Nos. 4,259,639 to Renirie, 4,231,027 to Mann et al., and 4,324,251 to Mann. These patents are also hereby incorporated by reference herein in their entirety. The theory underlying the use of internal impedance as a EOL warning indicator is that at low current drains typical of implantable medical devices, plots of resistance versus time give more warning than plots of terminal voltage over time. If voltage characteristics for different current drains are considered, the knees in the impedance curve are observed to have a fairly wide variation, meaning that the voltage at which the knee might appear is similarly subject to substantial variation as a function not only of the particular battery being used but also of the current being drawn by the pacemaker circuitry at a given time. On the other hand, plots of resistance indicate that the knee varies over a smaller range of values of internal resistance. Since the current drain may vary drastically with different electrode loads, the variation in voltage may be twice as great as the variation in internal resistance. Monitoring the internal resistance thus provides a more direct indication of the depth of discharge of the battery, whereas monitoring the output voltage gives a much less direct indication, reflecting not only the depth of discharge but also the current drain.
Provision of an ERI is not the only reason for monitoring a battery's depth of discharge. Another reason is that, as noted above, control of output pulse energy levels may also require information about the battery's output. In the prior art, it has been common to provide charge amplitude control circuitry for charging the output capacitors to a selected voltage. As previously noted, however, the charge amplitude control circuitry itself may consume battery power. In recognition of this, some pacemakers in the prior art have been designed such that when the battery reaches a particular level of depletion, the charge amplitude control circuitry is disabled, in order to minimize power consumption. This is done, for example, in the Cosmos II pacemaker manufactured and commercially available from Intermedics, Inc., Freeport, Tex. After the battery has depleted to such a level that the control circuitry is disabled, the capacitors are charged in an unregulated manner, i.e., without monitoring of the charging voltage. A worst-cause battery voltage is assumed, so that the capacitors are charged to at least some minimum level.
While this prior art arrangement eliminates current drain due to output regulation circuitry near the end of the battery's life, it has the disadvantage of reducing the pacemaker's ability to accurately control output pulse energy levels after the battery has reached a given level of depletion. In addition, such an arrangement may not be optimal in terms of device longevity. Thus, the inventor believes that there remains to be a need for improved arrangements for selecting and controlling output pulse energy levels in implanted devices.